CN111742215A

CN111742215A - Metal-enzyme interlayers

Info

Publication number: CN111742215A
Application number: CN201880089900.4A
Authority: CN
Inventors: 德温·艾达里安; 萨姆森·陈; 穆罕默德·M·吉拉尼; 阿克塞尔·谢勒; 希奥玛拉·林奈特·马德罗; 理查德·史密斯
Original assignee: California Institute of Technology CalTech
Current assignee: California Institute of Technology CalTech
Priority date: 2018-04-09
Filing date: 2018-11-30
Publication date: 2020-10-02
Anticipated expiration: 2038-11-30
Also published as: EP4235165A2; EP3775861A2; CN111742215B; US11109784B2; US20190307378A1; WO2019209377A2; EP4235165C0; EP3775861B1; US20220071529A1; EP4235165A3; WO2019209377A3; EP4235165B1; EP3775861A4

Abstract

The measurement of the target analyte is performed with an enzyme-based sensor. The enzymatic hydrogel is protected by a porous layer of a metallic material. The pores are small enough to prevent degradation of the enzyme layer by the organism's immune system, but large enough to allow transfer of molecules involved in the electrochemical reaction, thereby allowing the enzyme to detect the target analyte.

Description

Metal-enzyme interlayers

Cross Reference to Related Applications

This application claims priority from U.S. provisional patent application No. 62/654,701 filed on 9.4.2018 and U.S. patent application No. 16/191,357 filed on 14.11.2018, the disclosures of which are incorporated herein by reference in their entireties.

Benefit statement

The invention was made with government support under approval number HR0011-15-2-0050 granted by DARPA. The government has certain rights in this invention.

Technical Field

The present disclosure relates to biosensing. More particularly, the present disclosure relates to metallo-enzyme interlayers (metal-enzymesandwich layers).

Brief Description of Drawings

The accompanying drawings, which are incorporated in and constitute a part of this specification, illustrate one or more embodiments of the disclosure and, together with the description of the example embodiments, serve to explain the principles and implementations of the disclosure.

Fig. 1-2 illustrate steps in a sensor manufacturing process.

Fig. 3 illustrates an electrode configuration.

Fig. 4 illustrates an image of irregular peaks.

Fig. 5 illustrates an exemplary periodic (periodic) nanopillar.

FIG. 6 illustrates an exemplary sensor having a protective metal top layer.

Fig. 7 illustrates exemplary data.

Fig. 8 illustrates an exemplary method of creating an electrochemical sensor on a flat substrate.

Fig. 9-10 illustrate different embodiments of sensing devices.

FIG. 11 illustrates an exemplary array of sensing devices.

SUMMARY

In a first aspect of the disclosure, a sensing device is described, the sensing device comprising: a working electrode on the substrate; wherein: the working electrode includes: a metal layer contacting the substrate; an enzyme layer on the metal layer; and a metal porous layer on the enzyme layer, and the working electrode, the metal porous layer, and the enzyme layer are configured to detect a target analyte by an electrochemical reaction.

Detailed description of the invention

Layers containing enzyme molecules are important for the electrochemical measurement of glucose, lactate and many other medically important metabolites in the body. As is known to those of ordinary skill in the art, during chemical reactions involving analyte molecules, enzymatic reactions can be used to selectively convert an analyte of interest into a byproduct, releasing an electroactive compound. Electrochemical measurement of the rate of production of electroactive compound can then provide concentration data for the analyte molecules. These measurements are generally not error-free, as the enzymatic reaction may not be completely selective for the material of interest. In other words, the occurrence of an enzymatic reaction may be accompanied by false positives. Another factor is that the reagents required for the reaction to occur should be present in sufficient quantities otherwise the electrochemical measurement will be limited by the amount of reagents and not be completely accurate. Furthermore, the conversion rate of the enzyme should be high enough to produce an electroactive compound that is proportional to the analyte molecules, otherwise the signal may be lost in noise.

In some embodiments, the electroactive compound can be an ion. However, in some embodiments, the electroactive compound can be a molecule other than an ion. For example, the product of interest (hydrogen peroxide) in a typical glucose oxidase reaction is not an ion. In the present disclosure, some embodiments may refer to ions as electroactive compounds, however, more generally, these embodiments may also be practiced without ions but with uncharged electroactive compounds. In the present disclosure, "electroactive compound" refers to an ion and an uncharged compound capable of generating a current through an electrode by a chemical reaction on the surface of the electrode.

If significant current can be collected from the electroactive products, accurate measurements of metabolic biomarkers are possible as long as the enzyme molecule remains active. Thus, enzymatic degradation must be considered and prevented or at least reduced to extend the lifetime of the sensor. The efficiency of the conversion of the analyte reaction into a measured current is characterized by the conversion efficiency, which can be optimized if the electroactive products are efficiently collected on the measurement electrode and are not lost to the surrounding environment by diffusion. The reaction rate should also be limited by the presence of the analyte of interest and the concentration gradient of the enzyme molecule by the reagent maintained by the appropriate turnover rate.

Since the 60's of the 20 th century, enzymatic reactions have been used to measure glucose and other analytes. The first electrochemical glucose detector measures the current on the working electrode, minus 0.6V from the counter electrode, in order to collect charge from the hydrogen peroxide generated by the enzymatic reaction between glucose and oxygen catalyzed by glucose oxidase (GOx). In this catalytic reaction, glucose is oxidized to hydrogen peroxide and D-glucono-lactone (D-glucono-lactone). The current at the metal working electrode for the conversion of hydrogen peroxide to water and oxygen was measured with a potentiostat (potentiostat). Over a wide concentration range, the current is proportional to the glucose concentration. Since the beginning of the 60's of the 20 th century, many other enzymes (both oxidases and dehydrogenases) have been identified to measure other analytes of interest. Such enzymes can selectively react with glucose, lactate, alcohol, urea, cholesterol, xanthines, and tens of other analytes of interest.

In enzyme-based electrochemical sensors, the efficiency of conversion of analyte molecules to a measurable current typically changes over time as the enzyme molecule is attacked, dissolved or loses its cofactor. These processes result in a degradation of the sensitivity and accuracy of the sensor over time (deteriorations). There are several failure modes of electrochemical sensors that rely on enzymatic reactions. For example, enzymes may deteriorate due to oxidation due to molecular compounds in their surroundings or due to reactive compounds generated by the sensor itself. Enzymes may also lose cofactors that can react efficiently with the analyte. Other necessary reagents for the enzymatic reaction, such as oxygen for the oxidase, may be depleted, or the reactive enzyme membrane may simply delaminate by degrading its adhesion to the substrate. These failure modes are often accelerated when the sensor is located in a living organism, since the immune system of the organism attacks the sensor by its physical reaction (foreign body response) to foreign substances and leads to enzyme degradation by oxidation. The intracorporeal implant may be encapsulated, by which the immune response creates a scar tissue diffusion barrier around the implanted device, prevents reagents and analytes from penetrating to the sensor electrodes, and isolates the implant from the rest of the organism. All of these effects ultimately limit the lifetime of the sensor, or the length of time the electrochemical sensor can provide useful and accurate information. Most sensors are limited to operating in challenging environments encountered in living organisms for several weeks.

The present disclosure describes new methods for avoiding conventional degradation mechanisms and extending the lifetime of enzyme-based electrochemical sensors and other sensors that rely on monitoring selective (binding) reactions on or near electrochemically active contacts (contacts). Typically, electrodes are used to form electrochemical sensors by using the geometry of needles or wires or by defining electrodes on a flat surface. Ideally, the electrode material should be inert to the surrounding environment and not suffer from rapid corrosion. Platinum, gold, tungsten, iridium and titanium are therefore preferred materials because they are biocompatible materials, do not deteriorate rapidly in the body, and do not produce adverse immune system reactions.

The impedance of these electrodes ultimately determines the current that can flow through the electrodes at a given voltage, and in order to maximize sensor sensitivity and signal-to-noise ratio, it is desirable that the electrodes minimize the circuit resistance. This impedance can be varied by varying the surface area of the electrodes. The increased surface area may be achieved by increasing the lateral dimension of the sensor or by fabricating a three-dimensional electrode surface with increased surface area for the same lateral dimension. For example, the surface area may be increased by corrugation (corrugation), by expanding the sponge-like geometry, or by making regular or irregular nano-pillars or saw-tooth like structures. Generally, electrodes with large surface area to volume ratios can provide low impedance contacts for electrochemical reactions.

In conventional electrochemical sensors that rely on monitoring ions generated by enzymatic reactions, the surface of a metal contact is coated with a thin adhesive layer containing an enzyme. This layer is often referred to as a "hydrogel" because it is composed of a relatively porous material that enables smaller molecular analytes (glucose, lactate, etc.) to diffuse through the layer to react with enzyme molecules immobilized within the molecular structure of the layer. The traditional chemistry used in glucose monitors is the glutaraldehyde system, a polymer that is cross-linked to form a chemically stable scaffold in which enzyme molecules can be immobilized. The polymer is porous enough to keep the enzyme molecules active and allow reagents to enter and products to leave the polymer. In commercial sensors, the glutaraldehyde-based layer containing the enzyme is typically several microns thick. Enzyme molecules generate ions that need to pass through the thickness of a few microns in order to be read at the electrode surface. Unfortunately, any residual glutaraldehyde is immunogenic and can accelerate the innate immune response. Thus, glutaraldehyde-based layers are typically covered with a protective layer consisting of polyurethane or pegylated with polyethylene glycol to reduce the immune response. These additional layers generally reduce the sensitivity of the enzyme sensor and are slowly attacked when they are attacked by naturally oxidizing substances (oxidizing species), which in turn leads to a relatively short sensor lifetime.

The present disclosure describes a method for extending the lifetime of an enzyme-based electrochemical sensor. The method uses a thin metal layer deposited on the top surface of a thin layer containing the active enzyme. For example, platinum may be deposited in a very porous geometry, commonly referred to as "platinum black". Other materials such as titanium, gold or tungsten may also be used to fabricate a metallic porous layer that allows the target analyte, reagents and byproducts to move in both directions through the porous electrode, thereby allowing access to the underlying enzyme hydrogel. The platinum (or other metal) layer may be deposited by a vacuum deposition process in which platinum is sputtered from a platinum target and deposited onto the sample. The sputter deposition process is capable of controlling both the thickness and microstructure of the platinum layer by adjusting deposition parameters, including voltage, power, and vacuum pressure during deposition. Thus, the platinum black layer can be used as a low resistance contact surface for the electrochemical electrode. Sputter deposition is not the only method of depositing metal, and an alternative method involves evaporation from a heating source. The source may be heated by resistive methods or by electron beam heating. In other embodiments, other deposition methods may be used, such as pulsed laser deposition. A key requirement of the deposition technique is that the local heating of the enzyme layer and the time required for deposition should be minimized. In fact, localized heating may degrade the enzyme layer. While material deposition may require corrosive solutions, it is also possible to deposit a metal layer from a plating bath by electroplating or electroless plating (electroplating).

The present disclosure describes the deposition of a platinum/enzyme/platinum interlayer or multilayers rather than depositing a single enzyme-containing active layer on top of a platinum electrode layer. The structure includes a top platinum layer (which is relatively porous to enable the analyte to reach the enzyme molecules in the intermediate layer) and a bottom platinum layer (which acts as an electrode). This approach brings several important advantages. The platinum layer shields the underlying enzyme layer from oxidizing substances that would attack the polymer through the process of an immune response. The pores in the top metal layer are nanoscale, so the porous platinum layer prevents the immune system from detecting and recognizing potential chemicals (chemistry), removing immunogenic behavior. In fact, the active components of the immune system are generally larger than the nanoscale pores. The top platinum layer optionally serves as another contact surface that traps escaping ions that would otherwise be lost and enables them to contribute to electrical measurements. This is possible if the top platinum layer is electrically connected to the bottom platinum layer by vias (vias) through the enzyme layer or by electrical pathways around the enzyme layer. The platinum layer also retains oxygen, which is generated when peroxide ions react at the metal surface to convert to water and dissolved oxygen. Oxygen is required for the subsequent reaction, which requires oxygen as a reagent for the enzyme-assisted conversion to proceed successfully. The distance between the bottom and top platinum layers can be reduced to control the loss of cofactor by diffusing into the solution surrounding the sensor. All these effects have been shown to improve the linearity of the sensor response to glucose.

In some cases, it may be preferable to electrically attach the top porous metal layer to the bottom electrode. This process can effectively double the available electrode area, which may be desirable where sensitivity is important. Attachment may be made in several ways. As shown in the cross-section of fig. 9, the enzyme-based sensing layer (905) can be made slightly smaller than the bottom electrode layer (910). In other words, the lateral dimension of the enzyme layer is slightly smaller, leaving some space on the sides. The top conductive porous layer (900) can then be made larger than the enzyme-based sensing layer. In this way, the top layer is electrically connected to the bottom electrode layer through conductive pathways around the edge of the enzyme layer (periphery). Alternatively, the enzyme-based sensing layer (920) may be fabricated with a plurality of holes (holes) or vias (902) through which the top porous layer (915) is electrically connected to the bottom electrode layer (925). These holes may be formed by photolithographic patterning techniques, and in some cases, these holes may be formed by adjusting the deposition parameters of the enzyme-based sensing layer so that it naturally contains pores. One advantage of this manufacturing technique is that the large number of electrically conductive attachment points increases both the reliability of the electrical connection and its mechanical reliability. Another way to form these conductive attachment points is to fabricate the bottom electrode layer (945) such that it has pillars (904). An enzyme-based sensing layer (935) may then be deposited on the bottom electrode layer (945), leaving the top surfaces of the pillars exposed. When a porous layer (930) is deposited on the structure, it contacts the top surface of the pillar, forming a plurality of conductive paths between the top and bottom electrodes.

These conductive attachment points themselves may increase the effective electrode area because each conductive attachment point increases some surface area. Thus, where increased electrode area is beneficial, the width of each conductive attachment point can be as small as 5-40 nanometers, which is a limitation of most practical fabrication techniques for this type of structure. The spacing between attachment points (spacing) can be as small as 100 nanometers. However, such attachment points can be fragile and are often complex to produce. Larger attachment points such as 1 to 25 microns, and larger spacings such as 1 to 25 microns, are preferred if such a high effective surface area is not necessary or may not be beneficial due to higher background currents. The shape of these attachment points is generally determined by the manufacturing technique and may have, for example, a circular or rectangular cross-section. The height of these attachment points is typically selected based on the desired thickness of the enzyme sensing layer.

It may be preferable in some embodiments to ensure that the top conductive porous layer is electrically disconnected from the bottom electrode layer. In fact, the top conductive porous layer itself may generate an interference current from interfering compounds generated outside the sensor. Additionally, the increased surface area may increase background current. In some cases, these interference and background currents may be large enough to offset the benefit of the increased surface area for collecting molecules from the enzyme sensing layer. There are a number of ways to maintain an electrical disconnect between the top porous layer and the bottom electrode layer. In fig. 10, one way of maintaining electrical isolation is shown. The enzyme sensing layer (1005) is fabricated to have a larger lateral dimension than the bottom electrode layer (1000). The top porous layer (1010) is fabricated with a larger lateral dimension than the enzyme sensing layer, completely encapsulating the enzyme layer while avoiding any electrical connection between the top and bottom layers.

In some embodiments, the enzyme-based sensing layer naturally contains pores when deposited, resulting in the top layer being electrically connected to the bottom layer. In these cases, it may be necessary to deposit additional insulating layers. For example, if an imperfect (imperfect) enzyme sensing layer (1020) containing pores is deposited on the bottom electrode (1025), a thin porous electrical insulator (1015) may be deposited on top of the enzyme sensing layer (1020) to separate the bottom electrode (1025) from the top porous conductive electrode (1030). Can use a plurality ofPore insulating materials, including, for example, polyurethane and Nafion^TMThey are commonly used in electrochemical sensors for this purpose. In some cases, inorganic insulators with pores, such as silicon nitride or silicon dioxide, can be fabricated to accomplish the same task. Typically, the material used for layer (1015) must be an electrical insulator and contain pores large enough to readily allow the analyte of interest to reach the enzyme sensing layer. The thickness of this layer is determined by a number of factors, including the thickness required to maintain adequate electrical insulation, the porosity of the layer, the nature of the deposition method of the top porous layer, and the roughness and thickness of the underlying enzyme sensing layer. Typically, the layer is made as thin as possible so that it does not interfere with the transport of analytes or other molecules into and out of the sensing layer. For example, inorganic insulators generally have excellent insulating properties. If an inorganic insulator is used, a layer between 10nm and 500nm may be sufficient to perform this function. If an organic insulator is used, a layer thickness of 100nm to 5 microns may be used.

As known to those of ordinary skill in the art, alternative methods of obtaining linear sensor performance require the use of an organic control layer, such as polyurethane, to control diffusion. The exact thickness of these layers and their adhesion chemistry are very dependent on the ambient humidity and temperature at which they are deposited. Variability is difficult to control and can introduce uncertainty into the measurements. The porous metal diffusion control layer described in the present disclosure provides a solution with improved ion collection efficiency and oxygen recycling, since no peroxide molecules are lost to the surrounding tissue, resulting in little decrease in the measured sensor sensitivity. Thus, the use of a porous metal layer reduces variability between different sensors, which in turn improves the accuracy and reliability of each sensor.

In addition to providing better linearity and oxygen insensitivity, the metal surface protective layer also protects the enzyme molecules from attack by hiding the sensing chemistry from the immune system, resulting in a highly efficient electrochemical sensor with a very thin enzyme layer that offers a long lifetime. The thickness of the enzyme layer determines the delay between the reaction of the sensor to the analyte molecule of interest and the observation of an electrical signal from the ions released at the contact surface. In the structure of the present disclosure, this delay can be shortened, limited by the speed of the enzyme reaction, determined by the turn-over rate, rather than by the diffusion rate of the ions in the polymer matrix.

Metal deposition by vacuum processing can introduce a number of potential drawbacks when thin and brittle organic materials are to be coated. However, by appropriate polymerization of the enzyme layer with glutaraldehyde and subsequent metallization using a platinum sputter deposition process, it is possible to deposit a metal layer on the enzyme without degrading the chemical properties of the enzyme too much. The metal/enzyme/metal interlayer may even be deposited on a photoresist lift-off mask and patterned by using a photolithographic method. This fabrication method leaves the active interlayer only on top of the electrode of interest and thus enables functionalization of the selected regions with micron accuracy. The lift-off chemistry, which consists of selectively dissolving the photoresist template layer with acetone to define a pattern of metal/enzyme/metal functionalized areas, did not significantly reduce the enzymatic activity of the glucose oxidase enzyme and allowed the sandwich to be defined on top of the electrochemical potentiostat detector. This may also apply to other enzymes, such as lactate oxidase or urease, and leads to the opportunity to define several different chemical sensors on the same substrate by a series of lithographic processes that deposit different enzyme layers onto the electrochemical working electrode. Thus, in some embodiments, the sensor may comprise more than one region, each region having a different enzyme layer, so that more than one target analyte is detected with the same sensor. Unlike peroxide or ammonia molecule-evasive sensor systems known in the art, the metal-based multilayers described in the present disclosure prevent these molecules from escaping, avoid cross-talk, and enable accurate measurement of several very close analytes.

In such conventional enzyme-based sensor systems, it is often very difficult to measure more than one analyte in a compact sensor due to cross-talk. In particular, many enzymes for different analytes produce the same reporter molecule (a reporter molecule is a molecule that produces a sense), and is produced as a result of one analyteFor example, an implantable sensor designed to measure both lactate and glucose may have two electrodes, one coated with lactate oxidase and one coated with glucose oxidase when glucose levels are high, hydrogen peroxide produced by glucose oxidase may diffuse to lactate oxidase electrodes, falsely raising lactate readings.these electrodes may typically need to be separated by a distance of 1mm or more to reduce crosstalk to acceptable levels, depending on the environment in which the sensor is implanted and on the accuracy required.A wireless implantable sensor may be less than 1mm, with electrodes as small as 200 × 200 microns, so when more than one analyte must be sensed, the required separation may limit the minimum device size²Can detect multiple analytes in a single device.

In some embodiments, the sensor may also include wireless electronics to transmit its measurements from within the human body to an external device. For example, the sensor may have dimensions of 1.2mm by 1.2mm, a fast response time of 0.1s, use low power of 5 microwatts, and take only 10 cents. The millimeter-scale sized sensor does not move relative to the surrounding cells, maintaining a constant impedance. Due to its small size, the sensor can also be implanted in many locations, which can be injected and produce little tissue irritation. Possible applications include monitoring of chronic diseases such as diabetes, renal failure, cardiovascular problems and cancer. Sensors may also be used in skin patches to monitor cortisol, alcohol, glucose, and applications such as sports medicine and post-operative monitoring. Power may be transmitted to the sensor by wireless transmission, for example using a coil antenna. Typical electrode configurations may be used, such as concentric (concentric) square working, reference and counter electrodes. For example, fig. 3 shows a configuration with a counter electrode (305), a reference electrode (310), and a working electrode (305).

In some embodiments, all three electrodes are fabricated using platinum. Platinum allows the use of lower voltages and more stable operation compared to known Ag/AgCl electrodes. Fig. 4 shows atomic force microscope images (405,410) of exemplary electrode structures with higher irregular peaks for high current (e.g., 500nA) sensors. Fig. 4 also shows lower peaks (415,420) for a low current (e.g., 5nA) sensor. Fig. 5 shows an exemplary nanopillar that may be fabricated on the surface of an electrode to increase its surface area and sensitivity. The nano-pillars may be of the same metal material to form a uniform electrode. In some embodiments, the sensor is fabricated by spin coating the enzyme rather than drop coating. Spin coating allows for thinner electrode layers. A thinner enzyme layer increases the collection efficiency of the sensor, since hydrogen peroxide must diffuse to the surface of the metal electrode to be measured. In some embodiments, the enzyme hydrogel layer is no thicker than 1 micron, and is typically between 2nm and 1 micron. The enzyme molecules are typically larger than 1-2 nanometers, so the smallest enzyme layer is typically the size of the enzyme molecule. The top electrode controls the transfer rate of diffusion of the target analyte through the pores in the metal to the enzyme layer as well as the transfer rate of the peroxide. The top metal layer may have holes (or pinholes) between 1nm and 1 micron. The metal layer forms an electrochemical capacitor, and thus its porous structure affects the electrical performance of the sensor. The bottom metal layer may be of a desired thickness as it provides structural support for the other layers.

In some embodiments, platinum provides the best choice for the porous layer because it matches the working and reference electrode materials, it is inert, and long lasting. In some embodiments, gold may be used because it is not very reactive and may generally be considered biocompatible. Gold can complicate measurements within an electronic battery because a galvanic voltage (gal) may exist between gold and a reference electrodeA sonic voltage) and an electric current. Tungsten may also be used because it is inert and generally biocompatible, but may also complicate measurements through galvanic interactions. Titanium is a biocompatible material and therefore titanium may also be used, although it tends to be oxidised but the oxide is electrically conductive and thus oxidation is not usually a problem but electrochemical interactions may be a problem. Palladium is a material used in many implants and can reasonably be considered biocompatible. Iridium is biocompatible but very expensive. Silicon has some electrical conductivity and it is reasonably biocompatible, however, it may create galvanic interactions. In some embodiments, the above metals or other metals, as well as alloys thereof, may be used. Semiconductors and conductive metal oxides such as, for example, TiO may also be used₂、WO₃、SnO₂、InO₂And the like. The deposition technique used may be vacuum techniques such as sputter deposition and vapor deposition, as well as electroplating or electroless plating.

The pores in the metal protective structure are large enough to let glucose, oxygen and water pass through, but small enough to avoid cells and macromolecules attacking the enzyme. Thus, in some embodiments, the pore size or diameter may range from 2 nanometers to 2 micrometers. In some embodiments, the pore size is below 200 nanometers, or below 20 nm. Controlling the porosity of the layer can control the amount of oxygen that can escape from the enzyme layer and the amount of glucose that can enter the enzyme layer. By considering the area of the pores as the open area of the sensor allowing the passage of the chemical substance and comparing it with the closed area of the metal around the pores, it is possible to define the porosity of the sensor. In some embodiments, the porosity may be between 0.001% and 50%. In some embodiments, the porosity is between 0.1% and 10%, and can be controlled by controlling deposition parameters. In other words, the open area of the pores is between 0.001% and 50% of the total area of the porous layer, or between 0.1% and 10% of the total area of the porous layer.

In some embodiments, the layer thicknesses of the three individual layers of the sandwich structure are as follows. The thickness of the bottom (contact) layer may be between 1nm and 1mm, since it is necessary to provide a conductive base. Pins may be used or a metal layer may be deposited on a silicon chip or polymer support. The thickness of the top (protective) layer may be between 1nm and 10 microns. In some embodiments, the top metal layer is thinner than 1 micron, or thinner than 100 nm. In some embodiments, the enzyme layer may be between 10nm and 50 microns. In some embodiments, the enzyme layer has a thickness of less than 10 microns, or less than 1 micron, or less than 400 nm.

Fig. 1 shows a metal layer (110) of the substrate (105) and the working electrode, for example a metal layer made of Pt by sputter deposition. The metal layer is patterned to create a higher surface area (115). The pattern may be irregular pillars such as in fig. 1, or regularly periodic pillars as in fig. 5. As shown in fig. 2, the enzyme layer (205) is spin coated on the patterned metal layer. For example, an enzymatic hydrogel may be used. For example, enzymes and BSA coatings may be used. A glutaraldehyde layer (210) may be deposited, for example by vacuum evaporation, to crosslink the enzyme hydrogel. For example, glutaraldehyde in an aqueous solution in an open container is inserted into a vacuum chamber, resulting in evaporation on the sample in the chamber. In some embodiments, the process is applied only to the working electrode, while in other embodiments it may also be applied to one or both of the other electrodes of fig. 3. In some embodiments, layer (210) in fig. 2 may include a glutaraldehyde base layer and a porous metal layer over the glutaraldehyde. In some embodiments, the layer (210) does not comprise glutaraldehyde, but only comprises a porous metal layer.

Fig. 6 shows an exemplary sensor with a protective metal top layer. The sensor includes a substrate (625), a first patterned metal layer (620), an enzyme hydrogel layer (615), and a second patterned metal layer that protects the enzyme hydrogel. The second layer includes a porous layer (610), the porous layer (610) having openings that allow a selected chemical to pass through while blocking other chemicals. The second layer may also have a patterned region (605) above it. For example, the top porous Pt layer may be 50nm thick, the enzyme layer may be 300nm thick, and the Pt underlayer may be 100nm thick.

Fig. 7 shows exemplary data comparing the superior performance of a sensor (710) with an uncoated enzyme layer with a sensor (705) with a protective porous layer as described in the present disclosure. In some embodiments, the sensor may comprise an enzyme such as glucose oxidase, lactate oxidase, xanthine oxidase, cholesterol oxidase, sarcosine oxidase, cortisol oxidase, urate oxidase, alcohol oxidase, glutathione oxidase, and nicotine oxidase. In some embodiments, the sensor may be attached to the skin with an adhesive patch, for example to monitor sweat or substances extracted by iontophoresis. The external sensor may be part of a fitness tracker with an external battery and communication module.

FIG. 8 illustrates an exemplary photolithography process: the sensor comprises three electrodes (805); photoresist is spun on the sensor (810); after exposure, portions of the photoresist are stripped (815), allowing deposition of multiple layers, such as enzymes and crosslinkers, and a porous metal protective layer; the remaining photoresist was completely stripped (820) and the three sensors now have an enzymatic hydrogel and a porous metal layer on top. For example, as can be seen in fig. 8 (825), three layers including a first metal layer, an enzyme layer, and a top porous metal layer may be deposited on the resist. In some embodiments, where the sensing device is fabricated by a photolithographic process, it is possible to protect the underlying enzyme molecules (i.e., exposure to acetone, etc.) during the photolithographic process using a metal deposited on top of the enzyme layer.

The present disclosure describes a structure consisting of a layer of reactive material on a substrate, the layer of reactive material being covered by a porous inorganic layer capable of controlling the flow of reagents and/or products into or out of the reactive material to control the rate of reaction. In some embodiments, the inorganic porous layer is platinum, gold, tungsten, iridium, titanium, carbon (including graphene or carbon nanotubes) or other biocompatible material with very low immunogenicity, and the reactive layer is enzymatic. In some embodiments, the inorganic layer is platinum or another catalytic metal capable of reducing active oxygen or oxidizing species before they can reach the sensitive reactive layer. In some embodiments, the sensor is capable of quantifying a specific analyte, the reactive layer is an enzyme in the oxidase family, and the porosity of the inorganic layer is controlled such that oxygen can freely enter the reactive layer, the analyte is limited, and other interfering or destructive substances (including immune cells) are rejected, improving the overall sensitivity, oxygen independence, and high analyte concentration performance of the system.

In some embodiments, the sensor is part of a power generation unit, the reactive layer is an enzyme, and interfering or damaging substances (including immune cells) are rejected outside the compartment. In some embodiments, the sensor includes a reactive material layer on a substrate, the reactive material layer being covered by a porous conductive layer that is capable of controlling the flow of reactants and/or products into or out of the reactive material and electrochemically converting a chemical within the reaction chamber into and out of an external circuit, thereby increasing the conversion efficiency of the chemical. In some embodiments, the substrate itself is electrically conductive, and periodic or incidental conductive attachment points are formed between the porous conductive layer and the substrate, thereby greatly increasing the conversion efficiency of the chemical species of interest. In some embodiments, the substrate and the porous conductive layer are individually comprised of platinum, PEDOT: PSS, or other biocompatible conductive material capable of electrochemical switching, the conductive attachment points are comprised of a biocompatible conductive material, and the reactive layer is enzymatic.

In some embodiments, the sensor is capable of quantifying a particular analyte, and the porosity of the inorganic layer is controlled such that the analyte and reactant can freely enter the reactive layer, and other interfering or destructive species (including immune cells) are rejected, allowing electrochemically active products to be converted with very low loss due to the reactive layer being almost completely encapsulated. In some embodiments, the sensor is part of a power generation unit, and the porosity of the inorganic layer is controlled such that reactants can freely enter the reactive layer, and other interfering or damaging substances (including immune cells) are rejected, allowing the electrochemically active products to generate energy with very low losses due to the reactive layer being almost completely encapsulated. In some embodiments, the biosensor includes an oxidase-based reactive layer. In some embodiments, the pores are sized to promote the flow of oxygen to the reactive layer, reject interfering and damaging substances (including immune cells) from entering the reactive layer, reduce the loss of hydrogen peroxide from the reactive layer, and optionally reduce the flow of analyte to the reactive layer relative to oxygen to improve the linearity of the biosensor at high analyte concentrations.

In some embodiments, the porous layer and the substrate are comprised of platinum, which is nearly non-immunogenic and capable of catalyzing the reduction of destructive active oxygen species. In some embodiments, the nearly complete encapsulation of the reactive layer allows nearly all of the hydrogen peroxide produced by the reactive layer to be captured, and most of the oxygen produced by the oxidation of the hydrogen peroxide to be reused by the reactive layer, improving the sensitivity and linearity of the biosensor.

In some embodiments, the porous conductive electrode is partially covered and filled with a reactive material such that conversion of the desired reaction product is maximized and the reactants for the reactive material regenerated by the conductive electrode are maximized. In some embodiments, the conductive electrode is part of the sensor and is comprised of platinum, the reactive material contains an oxidase, the desired reaction product is the analyte of interest, the regenerated reactant is oxygen, thereby increasing the linearity of the sensor, the linearity of the sensor at high analyte concentrations is increased, and the overall sensitivity is increased. In some embodiments, the electrode is part of an energy harvesting system.

In some embodiments, the enzyme layer is deposited by dipping, ink jet printing, spin coating, and the top protective metal is deposited by vacuum deposition techniques to intentionally contain pinholes or other microfabricated conduits for passing products and reagents through the metal film. In some embodiments, the metal layer of the electrode comprises an irregular jagged peak or nanopillar array. In some embodiments, the device may be implanted internally. In some embodiments, the three electrodes have concentric square shapes as in fig. 3.

In some embodiments, the sensing device may include only two electrodes, rather than three electrodes. In these embodiments, the two electrodes each comprise multiple layers comprising a protective porous metal layer as described in the present disclosure. In other embodiments, the sensing device may include a single electrode, rather than three or two electrodes. For example, an electrode system uses the body of the organism being measured as ground potential (ground) and makes measurements using a single working electrode.

In some embodiments, the top porous layer is not necessarily metallic, but may be conductive in order to enhance the collection of any electroactive species and may be catalytically active against any destructive chemical species, such as reactive oxygen species produced by the immune system. For example, the top porous conductive layer may be comprised of PEDOT: PSS or polypyrrole. As explained with reference to fig. 10, in some embodiments, the metal porous layer is electrically disconnected from the metal layer.

In some embodiments, the porosity of the top layer is intrinsic to the deposition process. For example, adjusting deposition parameters during electron beam evaporation or sputtering of a metal film can be used to control porosity. In some embodiments, the porosity of the top layer may be more intentionally patterned. For example, electron beam lithography (lithograph) or photolithography may be used to select the precise location and size of the holes. For example, a mask containing a pattern of holes may be deposited on top of a non-porous top layer and subsequently used to etch controlled holes into the top layer. In this way, the porosity of the top layer can be better controlled than would be possible when the membrane is naturally deposited with pores.

In some embodiments, the bottom sensing layer need not be metal, but rather a conductive electrode material. For example, carbon, polypyrrole and PEDOT are commonly used as non-metallic electrode materials. In some embodiments, the sensing devices may be arranged in an array. Thus, the biosensor may comprise an array of sensing devices as described in the present disclosure, wherein the spacing between the sensing devices of the array of sensing devices is configured to avoid cross-talk due to leakage of intermediate reporter molecules generated during the sensing process. Fig. 11 illustrates an exemplary array of sensing devices, where each device (1105) is as described above in this disclosure.

The examples set forth above are provided to those of ordinary skill in the art as a complete disclosure and description of how to make and use embodiments of the present disclosure, and are not intended to limit the scope of what one or more inventors regard as their disclosure.

Modifications of the above-described modes for carrying out the methods and systems disclosed herein that are obvious to those skilled in the art are intended to be within the scope of the following claims. All patents and publications mentioned in the specification are indicative of the levels of skill of those skilled in the art to which the disclosure pertains. All references cited in this disclosure are incorporated by reference to the same extent as if each reference had been individually incorporated by reference in its entirety.

It is to be understood that this disclosure is not limited to a particular method or system, which can, of course, vary. It is also to be understood that the terminology used herein is for the purpose of describing particular embodiments only and is not intended to be limiting. As used in this specification and the appended claims, the singular forms "a," "an," and "the" include plural referents unless the content clearly dictates otherwise. The term "more than one" includes two or more of the indicated item unless the content clearly dictates otherwise. Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this disclosure belongs.

Claims

1. A sensing device, the sensing device comprising:

a working electrode on the substrate;

wherein:

the working electrode includes:

a metal layer contacting the substrate;

an enzyme layer on the metal layer; and

a porous layer of a metal on the enzyme layer, and

the working electrode, the metal porous layer, and the enzyme layer are configured to detect a target analyte by an electrochemical reaction.

2. The sensing device of claim 1, wherein the metal layer comprises an array of nanopillars.

3. The sensing device of claim 1, wherein the metal layer comprises irregular saw-tooth like peaks.

4. The sensing device of any one of claims 1-2, wherein the metallic porous layer comprises an array of nanopillars.

5. The sensing device of claim 3, wherein the metallic porous layer comprises irregular jagged peaks.

6. The sensing device of any one of claims 1-5, wherein the metallic porous layer comprises an array of pores having a diameter larger than a diameter of molecules participating in the electrochemical reaction.

7. The sensing device of any one of claims 1-5, wherein the metallic porous layer comprises an array of pores having a diameter between 2nm and 2 microns.

8. The sensing device of any one of claims 1-5, wherein the metallic porous layer comprises an array of pores having a diameter of less than 200 nm.

9. The sensing device of any one of claims 1-5, wherein the metallic porous layer comprises an array of pores having a diameter of less than 20 nm.

10. The sensing device of any one of claims 6-9, wherein the array of pores has a total area of between 0.001% and 50% of the total area of the metallic porous layer.

11. The sensing device of any one of claims 6-9, wherein the array of pores has a total area of between 0.1% and 10% of the total area of the metallic porous layer.

12. The sensing device of any one of claims 1-11, wherein the metallic porous layer is made of a material selected from the group consisting of: pt, Au, W, Ti, Pd, Ir, Si, TiO₂、WO₃、SnO₂Graphene and InO₂。

13. The sensing device of any one of claims 1-12, wherein the metal layer has a thickness between 1nm and 1 mm.

14. The sensing device of any one of claims 1-13, wherein the metallic porous layer has a thickness between 1nm and 10 microns.

15. The sensing device of any one of claims 1-14, wherein the enzyme layer has a thickness of less than 10 microns.

16. The sensing device of any one of claims 1-14, wherein the enzyme layer has a thickness of less than 1 micron.

17. The sensing device of any one of claims 1-14, wherein the enzyme layer has a thickness of less than 400 nm.

18. The sensing device of any one of claims 1-17, wherein the enzyme layer comprises an oxidase and a hydrogel.

19. The sensing device of any one of claims 1-18, wherein the enzyme layer comprises an oxidase selected from the group consisting of: glucose oxidase, lactate oxidase, xanthine oxidase, cholesterol oxidase, sarcosine oxidase, cortisol oxidase, urate oxidase, alcohol oxidase, glutathione oxidase, and nicotine oxidase.

20. The sensing device of any one of claims 1-19, wherein the sensing device is configured to be implanted inside a living organism or attached to the outside of the skin of the living organism.

21. The sensing device of any one of claims 1-20, further comprising:

a counter electrode on the substrate; and

a reference electrode on the substrate and having a plurality of electrodes,

wherein:

the counter electrode includes:

a metal layer contacting the substrate;

an enzyme layer on the metal layer; and

a porous layer of a metal on the enzyme layer, and

the reference electrode includes:

a metal layer contacting the substrate;

an enzyme layer on the metal layer; and

a porous layer of metal on the enzyme layer.

22. The sensing device of claim 21, wherein the working electrode, the counter electrode, and the reference electrode have concentric square shapes.

23. The sensing device of any one of claims 1-22, wherein the target analyte is glucose and the enzyme layer comprises glucose oxidase.

24. The sensing device of any one of claims 1-23, wherein the metal porous layer is electrically connected to the metal layer.

25. A sensing apparatus according to claim 24, wherein:

the enzyme layer includes more than one hole, and

the holes are filled with the metal porous layer, thereby forming an electrical connection between the metal porous layer and the metal layer.

26. A sensing apparatus according to claim 24, wherein:

the enzyme layer is laterally smaller than the metal layer, and

the side of the enzyme layer is filled with the metal porous layer, thereby forming an electrical connection between the metal porous layer and the metal layer.

27. A sensing apparatus according to claim 24, wherein:

the metal layer includes a pillar, and

the spaces between the pillars are filled with the enzyme layer, thereby forming an electrical connection between the metal porous layer and the top surfaces of the pillars of the metal layer.

28. The sensing device of any one of claims 1-23, wherein the metal porous layer is electrically disconnected from the metal layer.

29. The sensing device of claim 28, wherein the enzyme layer is laterally longer than the metal layer, and the enzyme layer covers sides of the metal layer, thereby preventing electrical contact between the metal layer and the porous metal layer.

30. The sensing device of claim 28, further comprising a porous insulating layer between the metal layer and the enzyme layer.

31. The sensing device of claim 30, wherein the porous insulating layer is polyurethane.

32. The sensing device of any one of claims 1-19, wherein the enzyme layer further comprises glutaraldehyde.

33. An array of sensing devices, each sensing device of the array being as described in any one of claims 1-32, wherein a spacing between sensing devices of the array of sensing devices is configured to avoid cross-talk due to leakage of intermediate reporter molecules.

34. The array of claim 33, wherein the pitch is between 10nm and 250 nm.